Implantable device for long-term delivery of drugs

ABSTRACT

A device for sustained delivery of a poorly water soluble drug is described. A drug reservoir within the device, when in operation, contains an aqueous suspension of the drug mixed with a suspension of an excipient that, in one embodiment, generates acidic groups for a sustained period of time to maintain a desired pH in the aqueous suspension that in turn provides a constant concentration of a soluble form of the drug.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.12/918,369, filed Aug. 19, 2010, now allowed, which is a U.S. NationalStage of International Patent Application No. PCT/US2010/027037, filedMar. 11, 2010, which claims the benefit of U.S. Provisional ApplicationNo. 61/159,742, filed Mar. 12, 2009, each of which is incorporatedherein by reference in its entirety.

TECHNICAL FIELD

The subject matter described herein relates to drug delivery, and inparticular, to an implantable drug-delivery device designed for deliveryof a therapeutic agent at a constant rate over an extended period.

BACKGROUND

For a variety of therapeutic agents, it would be desirable to deliver anactive pharmaceutical ingredient (e.g., an agent or a drug) into thebloodstream from a subcutaneously implanted device in a subject at asubstantially constant rate over a sustained period of up to severalmonths. For selected drugs, this delivery pattern can providesubstantial clinical benefits to patients and address important unmetmedical needs.

In general, there are two challenges that must be overcome inimplementing an effective, long-term drug-delivery device of this type.First, the amount of drug delivered by the implanted device must besufficient to provide the desired therapeutic effect and besubstantially constant over time; that is, the release profileapproximates zero-order kinetics, so that the treated individualreceives a substantially constant therapeutic dose over a specific timeperiod without dose spiking or periods of sub-therapeutic delivery.Secondly, the device should be capable of holding an amount of drugsufficient for releasing a therapeutic dose of compound over an extendedperiod, e.g., 1-6 months, with a size and shape suitable forimplantation in a selected anatomical site. For example, a deviceintended to be implanted in a subcutaneous site preferably has anelongate shape and a cross-sectional depth of less than about 5-6 mm soas be accommodated in the limited depth of the subcutaneous space andnot to produce an unsuitably large bulge in the skin above theimplantation site. The device would preferably need to less than about50 mm in overall length so that normal movement would not cause thedevice to erode the surrounding tissues, particularly at the ends of thedevice where, during normal movement, bending of the device relative tothe plane of tissue may occur resulting in rupture of the device throughthe skin surface. Given these constrains, the maximum practical volumeof the drug reservoir of a subcutaneously implanted device is generallyconsidered to be in the range of 500 microliter (μL), assuming thatsubstantially all of the volume enclosed by the device walls isavailable to serve as a drug reservoir.

A preferred shape for a subcutaneously implantable device iscylindrical. Cylindrical devices may be implanted by placing the devicein an implanter tool or trocar, an open-ended, pointed cannula with aninner diameter slightly larger than the outer diameter of the device.The trocar, loaded with the device, is inserted, through a smallincision, and tunneled under the skin distally from the entry point. Thedevice is positioned by retracting the trocar shaft mechanically, or byremoving the trocar while placing pressure on the end of the deviceusing a rod or plunger passed through the opposite end of the trocarshaft, leaving the device in place under the skin.

The art describes implantable drug delivery devices. For example,implantable osmotic pumps are known (e.g., U.S. Pat. No. 5,728,396).These drug delivery pumps suffer from a lack of sufficient internalvolume for extended delivery of low-water soluble drugs at a therapeuticrate because the osmotic engine occupies as much as 50% of the availableinternal volume. These drug delivery pumps are also prone to clogging ofthe exit port by precipitation of drug held in solution within thereservoir which can lead to rapid shut down, possible device ruptureand/or dose dumping.

The foregoing examples of the related art and limitations relatedtherewith are intended to be illustrative and not exclusive. Otherlimitations of the related art will become apparent to those of skill inthe art upon a reading of the specification and a study of the drawings.

BRIEF SUMMARY

The following aspects and embodiments thereof described and illustratedbelow are meant to be exemplary and illustrative, not limiting in scope.

In one aspect, a drug delivery device is provided. The device comprisesa non-erodible, non-porous housing member defining a reservoir, thehousing member having first and second opposing ends. A porous partitionis positioned in the first end of the housing member, and containedwithin the reservoir is a drug formulation comprised of a sparinglysoluble drug and a solubility-modifying excipients. Thesolubility-modifying excipient is effective to provide a concentrationof the drug, in an aqueous suspension when the drug formulation ishydrated, sufficient to provide release of a therapeutic dose of thedrug from the device over a period of more than one month, as solubledrug diffuses out of the device across the partition.

In one embodiment, the housing member is water impermeable.

In another embodiment, the housing member is a metal.

In one embodiment, the porous partition is selected from a porouspolymer membrane, a sintered metallic membrane, and a ceramic membrane.

In one embodiment, the solubility-modifying excipient is abiocompatible, bioerodible polymer. In an exemplary embodiment, thepolymer is selected from polylactides, polyglycolides, copolymersthereof and polyethyleneglycol. In a preferred embodiment, the polymeris a co-polymer of polylactic acid and polyglycolic acid monomericunits, wherein the polylactic acid content is between about 50% to 100%.

In one embodiment, the drug is a neuroleptic agent. In exemplaryembodiments, the neuroleptic agent is risperidone, 9-hydroxyrisperidoneor a pharmaceutically acceptable salt thereof. In other embodiments, theneuroleptic agent is olanzapine, paliperidone, asenapine, haloperidol oraripiprazole or a pharmaceutically acceptable salt thereof.

In one embodiment, the total amount of neuroleptic agent loaded is thereservoir is greater than 100 mg.

In yet another embodiment, the drug is buprenorphine.

In still another embodiment, the drug is present in soluble andinsoluble forms in a total amount of greater than 100 mg/mL.

In one embodiment, the soluble fraction of drug is less than 1% of thetotal.

In another aspect, a method for delivering a sparingly soluble drug froman aqueous suspension into an environment of use is provided. The methodcomprises formulating the drug with an excipient effective to maintain asubstantially constant concentration of a soluble form of the drug inthe aqueous suspension to sustain release of a therapeutic amount of thedrug into an environment of use for at least 30 days.

In another aspect, a method of treating a patient is provided. Themethod comprises providing a device as described herein, and implantingthe device in the patient. In one embodiment, the method is for treatinga patient suffering from a psychotic disorder, wherein the drug is aneuroleptic agent and is delivered at a substantially constant releaserate for between 1-6 months.

In one embodiment, the device is implanted subcutaneously.

In another aspect, an implantable device for use in releasing atherapeutic agent at an implantation site in a subject, at asubstantially constant release rate over a selected time period betweenabout 1-6 months is provided. The device comprises, in operativecondition, an housing that defines an interior chamber, the housingformed of a non-erodible, non-porous material; a partition affixed to anend of the housing, the partition comprising a plurality of pores whichallow therapeutic agent in a soluble form, but not in an insoluble form,to diffuse out of the chamber into an external medium, and containedwithin the chamber, in operative condition when the chamber is hydrated,an aqueous suspension composed of a mixture of a poorly water solubletherapeutic agent having a soluble agent form and an insoluble agentform and a solubility-modifying excipient. The solubility-modifyingexcipient acts to produce a concentration of the soluble agent formsufficient to provide a therapeutic dose of the drug over a selectedtime period, as the soluble agent form diffuses out of the device acrossthe partition by means of a concentration gradient of soluble agent formin the chamber and in the environment of use.

In one embodiment, the device is cylindrical in shape, and its outersurface is devoid of pits, ridges or pores. One or both of the circularcylinder ends is fitted with the porous partition, and the overall outersurface area of the device is less than 7 cm².

In another aspect, an implantable device for use in releasing atherapeutic agent at an implantation site in a subject, at asubstantially constant release rate over a selected time period betweenabout 1-6 months, is provided. The device comprises an interiorreservoir defined by a housing member, and contained within thereservoir a therapeutic agent. The device also comprises a partitionseparating the reservoir from an external medium, the partitioncontaining multiple pores which allow therapeutic agent in soluble form,but not in insoluble form, to diffuse out of the reservoir into theexternal medium, the pores, in one embodiment, occupying less than 1.7cm² of the total surface of the chamber. Contained within the reservoir,in operative condition when the reservoir is hydrated, is an aqueoussuspension comprised of a mixture of the therapeutic agent and asolubility-modifying excipients. The therapeutic agent has a watersolubility at neutral pH such that the concentration of the agent in thewater phase of the suspension is insufficient to drive outward flux of atherapeutic dose of the agent. The excipients acts to produce aconcentration of a soluble form of the agent sufficient to provide atherapeutic dose over the selected time period, as soluble agentdiffuses out of the device across the partition by means of aconcentration gradient of soluble agent across the partition.

In one embodiment, the pH of said suspension is adjusted to be in therange of 2.5 to 6.8 and to control the equilibrium water solubility ofthe agent.

In another embodiment, the solubility-modifying agent is a polymermixture selected to erode at a rate that corresponds to the intendedperiod of operation of said device. In one embodiment, the polymer is apolylactide, polyglycolide co-polymer particle of a particle size,molecular weight, acid end-group concentration, monomer ratio, inherentviscosity and/or porosity as to erode and produce lactic acid andglycolic acid continuously for the intended period of device operation.

In addition to the exemplary aspects and embodiments described above,further aspects and embodiments will become apparent by reference to thedrawings and by study of the following descriptions.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a graphic representation of the theoretical relationshipbetween the release kinetics of an agent (rate over time and kineticorder) and the maintenance of a concentration gradient. Concentration ofthe agent is maintained approximately constant in an internal reservoirby balancing the flux rate and dissolution rates of the agent;

FIG. 2 is an illustration of one embodiment of an implantable drugdelivery device prepared for operation according to the principlesdescribed herein;

FIGS. 3A-3B illustrate another embodiment of an implantable device, withthe device shown in exploded view (FIG. 3A) and in assembled view;

FIGS. 4A-4B illustrate another embodiment of an implantable device inexploded view (FIG. 4A) and in assembled view (FIG. 4B), and an approachfor inserting a porous partition in one or both ends of the device;

FIGS. 5A-5B illustrate another embodiment of an implantable device inaccord with the teachings herein, the device shown in exploded view(FIG. 5A) and in assembled view (FIG. 5B);

FIG. 6 is a projection illustrating the approximate relationship betweenpH and the aqueous solubility of risperidone base;

FIGS. 7A-7B are graphs showing cumulative release of risperidone invitro, in mg (FIG. 7A), and the release rate of risperidone, in mg/day(FIG. 7B), as a function of time, in days, from devices having a porouspartition separating the drug formulation from the environment of use,the porous partition in the devices had a molecular weight cut-off(MWCO) of 100 KD (triangles) or 3 KD (squares) or was a 0.45 μmpolyvinyldenefluoride membrane filter (Durapore®; circles);

FIG. 8 is a graph of the cumulative release of risperidone, as a percentof total amount of risperidone in the device, as a function of time, indays, where the risperidone in the device was an aqueous solutioncomprising 5 mg (diamonds), 10 mg (squares) or 15 mg (triangles)poly-lactic-poly-glycolic acid (PLGA);

FIG. 9 is a bar graph showing the dose-response relationship betweenrisperidone release rate, in mg/day, and amount (5 mg, 10 mg, 15 mg) ofa solubility-modifying agent (85:15 PLGA) added to the aqueous drugsuspension loaded into the device;

FIG. 10 is a graph of risperidone release rate, in mg/day, as a functionof time, in days, for two devices in accord with the teachings herein,where one device included a mixture of two solubility-modifying agents(50:50 PLGA and 85:15 PLGA) in the drug suspension loaded into thedevice (closed circles) and the other device was filled with a drugsuspension lacking a solubility-modifying agent (open circles);

FIG. 11 is a graph of risperidone release rate, in mg./day, as afunction of time, in days, where the series of curves illustrates theinfluence of the solubility-modifying agent PLGA with different lacticacid:glycoloic acid ratios on the in vitro release of risperidone;

FIG. 12A is a graph showing the in vitro release rate, in mg/day, as afunction of time, in days, for two groups (n=3) of delivery devicesfitted at both ends with a porous partition and filled with an aqueoussuspension of risperidone with a solubilizing excipient (squares) orwithout the excipient (triangles); and

FIG. 12B is a graph showing the in vivo plasma concentration ofrisperidone and its active metabolite (9-hydroxyrisperidone), in ng/mL,as a function of time, in days, after subcutaneous implantation in ratsof the devices of FIG. 12A, where the drug formulation comprised(squares) or lacked (triangles) a solubility-modifying agent.

DETAILED DESCRIPTION I. Definitions

The term “aqueous suspension” refers to a solid material dispersed in aliquid solution that is substantially water.

The terms “therapeutic agent,” “drug,” “active pharmaceuticalingredient” or “API” refer to a biological or chemical agent used in thetreatment of a disease or disorder, or to treat or alleviate symptomsassociated with a disease or disorder.

The term “silicon nanopore membrane” means a membrane that has beenfabricated, for example, using micromachining techniques borrowed fromthe microelectronics industry.

The term “precipitate” or “insoluble form” refer to a solid substancethat separates from a solution.

A therapeutic agent is “relatively insoluble in water” or “sparinglysoluble in water” if its equilibrium solubility in water, measured atroom temperature, is less than about 1×10⁻³ M.

The term “solubility-modifying agent” refers to an excipient added tothe drug suspension formulation and which acts to increase the aqueoussolubility of the drug.

The term “substantially zero-order kinetics” means that over a medicallyacceptable percentage of the dose of a therapeutic agent provided in adrug delivery device, the rate of release of the agent is approximatelyconstant.

II. Drug Delivery Device

In a first aspect, a device for use in releasing a therapeutic agentinto an environment of use is provided. The therapeutic agent ispreferably one that is sparingly soluble in water. The device, forreasons described hereinbelow, is capable of providing a substantiallyconstant release rate of the drug over an extended time period. Thedevice is particularly well suited to deliver agents which are sparinglysoluble in water at neutral pH but become more water soluble at acidicpH. The device also allows delivery of agents which although slightlywater-soluble do not generate the concentration gradient sufficient todrive outward flux of a therapeutic dose of the agent. The device is ofa size, shape and surface properties to be suitable for subcutaneousimplantation and explanation. The device includes an interior chamberseparated from an external medium in an environment of use by a porouspartition, and contained within the chamber when the device is in use,is an aqueous suspension of the drug mixed with an excipient thatgenerates acidic equivalents. The appearance of these acidic moietieseffectively lowers the pH within the reservoir, and, in turn, improvesthe solubility of the drug. In one general embodiment, the excipient isa degradable polymer, and the continuous hydrolysis of the polymerprovides substantially constant release rate of the agent by maintaininga constant solution concentration of the agent in the chamber over amajor portion of the extended period of time, and thus providing asufficient concentration gradient between the chamber and the externalmedium to produce outward diffusion of the agent across the porouspartition at a rate that provides a therapeutic level of agent over theintended period of device operation.

FIG. 1 illustrates the relationship between the percent of total drugloaded into a drug delivery device (including both soluble and insolubleforms) as a function of time. As shown, as long as the concentrationgradient is maintained, the output rate is constant (i.e., zero-order).The insert in FIG. 1 presents Fick's law of diffusion and illustratesthat the flux (i.e., the outward movement of the agent) is directlyproportional to the concentration of the agent.

A. Drug Delivery Device Components and Assembly

Turning now to FIGS. 2-5 several exemplary embodiments of an implantabledrug delivery device for continuous delivery of a drug for an extendedperiod of time are illustrated. With initial reference to FIG. 2, adevice 10 intended for implantation in an anatomical compartment of asubject, such as under the skin or in the peritoneal cavity, isillustrated. The device is comprised of a housing member 12 that definesan internal compartment or reservoir 14. Contained within the reservoiris a drug formulation, described below. Housing member 12 has first andsecond ends, 16, 18. First end 16, in the device embodiment illustratedin FIG. 2, is sealed with a fluid-tight end-cap 20. Opposing second end18 is fitted with a porous partition 22. It will be appreciated that aporous partition can be positioned at one or both ends of the housingmember. As used herein, the terms “porous membrane” and “porouspartition” intend a structural member that has a plurality of pores inthe micrometer (μm) range, preferably in the 0.1-100 μm range. That is,the porous partition permits passage of the drug in soluble form fromthe drug formulation contained within the reservoir. The porouspartition can also permit passage of a solubility-modifying excipientthat is part of the drug formulation, discussed below, in its solubleform. The porous partition in a preferred embodiment retains the drugand/or the solubility-modifying excipient in their insoluble forms. Thatis, drug and/or solubility-modifying excipient in insoluble formpreferably do not pass through the pores of the porous partition.

Housing member 12 is preferably formed from a biocompatible material,and a skilled artisan will readily identify suitable materials,including but not limited to metals, such as titanium, and polymers. Apreferred shape of the housing member 12 is cylindrical. Skilledartisans, however, will understand that alternative geometrical shapesmay also be suitable, and these shapes are contemplated. In oneembodiment, a tubular section of the cylindrical housing member has anon-porous, smooth, and/or liquid impermeable outer surface. Forexample, a tubular section of housing member 12 formed of titanium andfinished on its outer surface, i.e., the surface of device 10 in contactwith the environment of use 25, has a satin, mirror, or other polish. Toinsure the safe and medically uncomplicated use, and removal, of thedevice, it is preferred that its outer surface be smooth, devoid ofpits, ridges and pores. Rough or porous surfaces may tend to provoke atissue response during the implantation period; tissue in-growth,attachment or adherence of matrix components or cellular processes orpseudopodia to pits or pores on the surface of the device can complicateremoval and possibly even tear the ingrown tissue upon removal, causingtrauma to the site. Smooth surfaced cylindrical devices, on the otherhand, provoke little or no tissue in-growth, are easily removed from theimplantation site by simply reopening the incision made duringimplantation and pushing the distal end of the device with one's fingersthrough the skin, forcing the device to emerge from the incision.Accordingly, in one embodiment, the housing member is non-porous, and inother embodiments is water-impermeable, non-erodible and/ornon-swelling.

Drug formulation 14 contained within the reservoir of the device ispreferably in the form of a continuous aqueous phase 26 which mayinclude a solubility-modifying excipient 24 in soluble form, insolubleform, or a mixture of soluble form and insoluble form. The drugformulation includes an active ingredient or drug, examples of which aregiven below, within the continuous aqueous phase. The drug is present inboth insoluble form, as represented by particle 28, and soluble form, asrepresented by 30. Equilibrium is established between the soluble andinsoluble forms of the drug within the drug formulation contained in thereservoir, as indicated by arrows 32. The relative mass of drug presentin soluble and insoluble forms depends on its intrinsic solubility inthe continuous aqueous phase, which may include a solubility-modifyingexcipients, the pH and the temperature.

As mentioned, device 10 is intended to be implanted into an anatomicalsite in a subject, and preferably in implanted at a subcutaneous site.Following implantation, the concentration of the drug in the externalmedium 25 is maintained at virtually zero at all times by the continuousmovement of interstitial fluid around the device. In operation, aconcentration gradient is established between the soluble form of thedrug or agent within the reservoir and the external medium. The gradientserves to drive diffusion of the soluble form of the agent from thereservoir through the porous partition into the external medium.Equilibrium is established between the insoluble fractions, generallyindicated at 35, and soluble fraction of the agent, generally indicatedat 37, within the reservoir and a released fraction, indicated at 39,which enters the external medium and is immediately washed away into thesystemic circulation. As the soluble form of the agent moves from thedevice reservoir, through the porous partition and into the externalmedium under the influence of the concentration gradient, insolubleagent in the formulation contained in the device reservoir dissolves,replacing that which has been released. In this manner a constantconcentration gradient is maintained and a constant outward flux ofagent is maintained according to Fick's law.

The diffusive flux, J, in mass/unit time, of a molecule through amembrane is defined by Fick's law as follows:

$\begin{matrix}{J = {D_{eff} \cdot A_{eff} \cdot \frac{\left( {C - C_{0}} \right)}{L}}} & (1)\end{matrix}$

wherein

-   -   D_(eff)=effective diffusion coefficient, cm²/sec    -   A_(eff)=effective Membrane area, cm²    -   C=concentration of diffusing species on one side of membrane,        mass/cm³    -   C₀=concentration of diffusion species on opposite side of        membrane    -   L=thickness of membrane, cm.        The effective membrane area is further given as follows:

A _(eff) =εA _(m)/100  (2)

wherein

-   -   ε=porosity of the membrane, %    -   A_(m)=physical membrane area, cm²

At constant temperature and for a typical type of porous partition (alsoreferred to herein as a microporous membrane), D_(eff) in Equation (1)is normally a function only of the molecular weight of the diffusingmolecule and in general D_(eff) decreases with increasing molecularweight. Therefore, Equation (1) indicates that for implant applicationswhere C₀ is approximately zero as it would be in vivo (sink conditions),providing a nearly constant soluble drug concentration will provide anearly constant drug diffusion rate from the implantable drug deliverydevice with specific selected properties (e.g., thickness and porosity)of membrane.

In the delivery devices, a chemical equilibrium between thewater-soluble and insoluble forms of the agent is established within thedrug reservoir. The porous partition separating the reservoir from theexternal medium is relatively thin (0.02-0.2 cm), highly porous and,preferably, hydrophilic. During operation the pores fill with theaqueous medium by capillary action; the pores thus provide a route forthe passage for the soluble fraction of the agent held in the reservoirto the external environment by a process of diffusion. The pores aresufficiently small as to prevent passage of insoluble drug particlesheld within the reservoir. In practice, the external medium isinterstitial fluid. Thus, after implantation, a concentration gradientis established across the partition; the concentration of the agent inthe interstitial fluid is virtually zero whereas the concentration ofthe agent in the reservoir equals the solubility of the agent within theinternal aqueous phase of the suspension. The flux of drug is regulatedby balancing the surface area of the partition (porosity) and theequilibrium solubility of the agent within in internal aqueous mediumaccording to Fick's second law of diffusion, wherein the desired outputrate in mg/day (J) is defined:

$\begin{matrix}{{J\left( {{mg}\text{/}{day}} \right)} = {\frac{mg}{8.64 \times 10^{4}} \times \frac{A_{M}}{100}}} & (3)\end{matrix}$

wherein A_(M) represents the surface area (i.e., porosity expressed aspercent of the area available for diffusion) of the partition.

The equilibrium aqueous concentration of the agent within the chamber(C_(i)) is adjusted to achieve the desired output rate according to thefollowing equation:

$\begin{matrix}{C_{i} = \frac{J \cdot T_{M}}{D}} & (4)\end{matrix}$

wherein D is the diffusion constant of the agent; and T_(M) is thethickness of the partition.

As mentioned above, the reservoir of the device contains an aqueoussuspension of a drug. The aqueous suspension within the reservoir isseparated from the external environment of use by the porous partitionwhich is freely permeable to water. When placed in the preferred in vivocompartment, the interstitial compartment, the carrier fluid within thereservoir would be in direct communication with interstitial fluid(which flows through the subcutaneous space) through the pores of theporous partition. Interstitial fluid is essentially an ultrafiltrate ofblood, containing the same ions, buffers and about 70% of the proteinpresent in blood, but excluding formed elements such as red blood cells.Similarly to blood, the pH of interstitial fluid is precisely regulatedto be about 7.4. Interstitial fluid continuously moves through thesubcutaneous space. All soluble components held within the reservoir ofa device incorporating a porous partition such as the one describedherein, including the carrier fluid itself, would be free to diffuse inboth directions across the partition. Rapid equilibration of ions andbuffers between the interstitial fluid and soluble components of thereservoir would ensue rapidly following implantation. For this reason,water miscible solvents such as dimethylsulfoxide (DMSO) or ethanol arenot preferred for use as the carrier fluid; such solvents would diffuseout of the device and be replaced by the inflow of interstitial fluid.Exchange of such a solvent carrier with the aqueous interstitial fluidwithin the reservoir would change the solubility of the loaded agent,adversely affecting its release rate and thus such water-misciblesolvents would be unsuitable. If a water-based carrier fluid is used,regardless of its initial composition and pH, equilibration of allsoluble species with those of the interstitial fluid would occur withinseveral hours following implantation. Accordingly, in one embodiment,the reservoir of the device does not include an organic solvent or doesnot include a water-miscible organic solvent. In another embodiment, thereservoir of the device, and the device itself, do not include anosmotic pump or engine.

FIGS. 3-5 further illustrate the implantable device by a discussion ofits assembly. FIG. 3A illustrates implantable drug delivery device 50 inexploded, cross-sectional view. A housing member 52 has a wall 54 withan external surface 56 and an internal surface 58. Housing member 52 ishollow, and defines an internal compartment 60. An exemplary housingmember is a titanium alloy tube, which are created by milling/lathingsolid titanium alloy material. The rod outer diameter (OD) is about 4.3mm and the length about 35 mm, in one embodiment. The internal reservoiris formed by drilling/milling out the center of each rod. For the deviceillustrated in FIG. 3A, a first end 62 of the housing member is open,and sized appropriately to permit additional device parts to be insertedinto and through the open end. As seen in FIG. 3A, first end 62 includesa first portion 62 a having a smaller outer diameter than the outerdiameter of the main portion 62 b of housing member 52. A lip ortransition 68 smoothly connects the first portion 62 a and the mainportion 62 b.

With continuing reference to FIG. 3A, a second end 64 has an annular rim66 extending inwards from the inner wall, and is referred to as a rimretention section. The rim retention section has, for example, an innerdiameter (ID) of 3.4 mm and a depth of a few hundredths of a mm. Thehousing member also includes a shoulder 70 directly adjacent theretaining rim. Shoulder 70 has a size wide enough to allow insertion ofthe device components parts, such as o-rings 72, 74 and a porouspartition 76, but small enough to capture and engage a press ring 78 forsealing the o-ring/porous partition/o-ring stack 80 (see FIG. 3B) intoplace at the second end 64 of the housing member. The ID of theremaining portion of the housing member, that is the portion betweenshoulder 70 and lip 68, is, in this embodiment, about 3-4.5 mm, morepreferably 3-5-4.0 mm, and preferably is 3.8 mm. The internal volume ofthe device is therefore approximately between 350-400 microliters, morepreferably 375-390 microliters, and preferably 380 microliters.

In one embodiment, the porous partition is a conventional porousmembrane with pore sizes selected to allow diffusion of the dissolvedagent but prevent insoluble material of leaving the device. In anotherembodiment, the partition may be made using silicon microfabricationtechniques which create assays of parallel channels each of which in itssmallest dimension is from 1.5 to 5 times the hydrodynamic diameter ofthe agent itself. When the channel width is properly tailored to themolecular dimensions of the agent, such nanopore membranes have beenshown to constrain the diffusion of the agent. Alternatively, the porouspartition may be selected from a variety of conventional crosslinkedpolymer membranes such as those used in ultrafiltration and dialysisapplications or crosslinked or polymerized gels such as polyacrylamide,agarose, alginate and the like. Sintered metal membranes or fritscomposed of titanium, titanium alloys or stainless steels or sinteredceramic membranes and metal screens typically used as in-line filtersalso may serve as the porous partition. Pore size and/or molecularweight cut-off of the membrane or gel would be selected to substantiallyretain the insoluble form of the drug within the device reservoir whileallowing diffusion of the agent across the membrane. In operation, therate of diffusion of the agent would be determined by the concentrationgradient of the agent across the membrane (i.e., the difference inconcentration of the soluble portion of the agent within the chamber and“sink conditions” external to the device) and the cross sectional area(porosity) of the partition (as exemplified in the Equations above). Inone embodiment, the total area occupied by the pores contained in thepartition is less than about 2 cm², more preferably less than about 1.7cm².

Device 50 is constructed by inserting into the reservoir the twoo-rings, 72, 74, with the porous partition 76 sandwiched between theo-rings, as indicated by arrow 90. The porous partition, in thisembodiment, is a stainless steel mesh screen or a sintered titaniumfrit. The o-rings are preferably made of silicone measuring 70 on theShore A scale, and preferably measure 0.091″ inner diameter by 0.026″cross section.

The partition press-ring 78 is then mechanically pressed into placeabove the o-ring 74, to apply pressure to the o-ring/partition/o-ringstack 80 and seal the o-rings against the inner wall of the housingmember, the porous partition and retention rim. In one embodiment, thepartition ring is an annular ring with a flat o-ring interface surface.When the o-ring is inserted into the housing member to its proper depth,it engages shoulder 70 of the housing member, forming a fluid-tight sealof the o-ring/partition/o-ring stack 80 with the retaining rim.

After the reservoir is filled with a desired drug formulation, and anexemplary filling procedure is described below, an end-cap 92 with adomed external surface 94, a press surface 96 to interface with the wallof the housing member and form a fluid-tight seal, is pressed onto end62, as indicated by arrow 98. FIG. 3B illustrates the device with itscomponent parts in place for implantation and operation.

FIGS. 4A-4B illustrate another embodiment of a drug delivery device, inexploded view (FIG. 4A) and fully assembled for use (FIG. 4B). Drugdelivery device 100 is comprised of a tubular housing 102 having a wall104 that defines an internal cavity 106. Housing 102 has opposing ends,106, 108. Housing 102 has a first outer diameter dimension at opposingends 106, 108 and a second outer diameter dimension in the portion ofthe housing between the opposing ends. Transition regions 110, 112connect the first and second outer dimensions of housing 102. A porouspartition, such a porous partition 114, is sized for contact with anannular rim, such as annular rim 115, at each end of the housing. Eachporous partition preferably has an outer diameter approximately equal tothe first outer diameter at each end 106, 108. A sealing member, such assealing member 116, is sized for engagement with each porous partitionabout the rim at each end 106, 108. An end cap, such as end cap 118, isdimensioned for a secure fit around the external surface of the housingat each end. In this embodiment, the end caps each have a centralopening, such as opening 120 in cap 118, to permit in-flow ofinterstitial fluid from the environment of use external to the deviceinto reservoir 106 and efflux of soluble drug from reservoir 106, acrossthe porous partition, and into the environment of use. FIG. 4B shows thedevice with the porous partitions, sealing members, and end caps inplace on housing member 102. As seen, the outer diameters of the endcaps are dimensioned for agreement with the a second outer diameterdimension in the portion of the housing between the opposing ends, sothat the overall outer dimension of the device when fully assembled isuniform.

FIGS. 5A-5B illustrate another embodiment of a drug delivery device, inexploded view (FIG. 5A) and fully assembled for use (FIG. 5B). Device130 comprises an annular encasement 132 have an external surface 134that is smooth, non-porous, and liquid impermeable. Encasement 132 ishollow, and its wall defines an internal cavity 136 which when assembledfor use contains a drug formulation. Each end of annular encasement,ends 138, 140, has a smaller outer diameter than a middle section 142 ofthe annular encasement. A frit, such as frits 144, 146, are dimensionedfor contact with a lip, 148, 150, at each end, frits 144, 146 having adiameter approximately equal to the outer diameter of ends 138, 140. Thefrits can be of ceramic, glass, or metal and are permeable to solubledrug in the drug formulation. A polymer membrane, such as membrane 152,abuts each frit and is in dimensional agreement with the frits. Thepolymer membrane can be porous or non-porous, so long as it is permeableto solubilized drug and fluid in the environment of use. An o-ring orsimilar sealing member, such as member 154 provides sealing engagementof the frit, membrane, and a press cap, such as press caps 156, 158.Together the frit, polymer membrane, and o-ring at each end of thedevice form a stack, such as stack 160, that is held in place at eachend by the press caps. Each press cap has a central opening, for fluidcommunication between the cavity of the device and the environment ofuse.

The internal reservoir of the drug delivery device is filled with a drugformulation. The formulation is comprised of a drug (also referred to asa therapeutic agent), a solubility-modifying agent, an aqueouscontinuous phase. The therapeutic agent used in the device may be asmall molecular weight drug, preferably having an equilibrium solubilityconstant of between about 1×10⁻³ M and 5×10⁻³ M at 37° C. in the aqueouscontinuous phase. In a preferred embodiment, the formulation in thedevice reservoir comprises an aqueous suspension of a drug that issparing soluble in water. The term ‘sparingly soluble’, as used herein,refers to a drug having a solubility measured in water at 37° C. of lessthan about 3 mg/ml at neutral pH, preferably of between about 0.001 to 3mg/mL, or 0.025 to 3 mg/mL, at neutral pH. More generally, sparinglysoluble drugs for use in the device described herein have a solubilityin water at 37° C. of less than about 3 mg/mL at neutral pH, and have anincreasing solubility as pH decreases, as discussed further below. Sincediffusion is driven by the concentration gradient established across theporous partition in the device, as discussed above, drugs with lowsolubility in water at physiological pH do not achieve a concentrationgradient sufficient to provide release of a therapeutic dose of thedrug.

In another embodiment, drugs for use with the device described hereinhave a therapeutic dose of less than about 3 mg/day, preferably of lessthan about 2.5 mg/day, and still more preferably of less than about 2mg/day or 1.5 mg/day.

Non-limiting examples of drugs includes neuroleptic agents, such asrisperidone and olanzapine. Neuroleptic agents are prescribed for thetreatment of psychotic disorders such as schizophrenia and bipolardisorder. The desired dose, or release rate, for neuroleptic agents is0.5-3 mg/day, more preferably 1-2 mg/day, for a period of between about1-3, 1-4, 1-5, or 1-6 months. Another class of exemplary therapeuticagents is the low-solubility opioid mixed agonist-antagonists, such asbuprenorphine, which is used to treat chronic pain, opioid addiction andalcoholism. Other sparingly-soluble drugs are paliperidone,aripiprazole, asenapine and haloperidol. These drugs, and especiallyrisperidone, olanzapine and buprenorphine, have low aqueous solubilityat physiological pHs, on the order of 1×10⁻³ M to 5×10⁻³M.

It is possible to adjust the pH of a drug suspension within a reservoirdevice that is separated by a porous partition from the external mediumto virtually any desired level and thus regulate the initial equilibriumsolubilities of drugs, such as risperidone. In the case of risperidone,solubility increases several fold as the pH is lowered to 6.0. Indeed byadjusting the internal pH of a risperidone suspension it is possible toraise the aqueous solubility from 0.4 mg/mL at pH 7.4 to about 100 mg/mLat pH 2.0, as shown in FIG. 6. For the device described herein, adequatedrug release rates would be possible at pH's below about 6.0. Inpractice, however, following implantation, exchange of buffers acrossthe partition (which is freely permeable both to water and smallmolecules), between the aqueous phase of the drug formulation held inthe reservoir of the device and the external interstitial fluid, wouldcause the internal pH of the drug formulation to quickly equilibrate to7.4, regardless of its initial pH. Accordingly, the drug formulationadditionally comprises an insoluble excipient capable of continuouslygenerating acid groups during operation of the device, effectivelydecreasing the pH of the drug formulation within the device reservoir tomaintain a pH that increases the aqueous solubility of the drug to alevel which provides a concentration sufficient to drive outwarddiffusion required to achieve a therapeutic in vivo release rate for aperiod of at least about 1 month, 2 months, or 3, 4, 5, or 6 months.

In one embodiment, the insoluble excipient, also referred to herein as asolubility-modifying excipient, is a polymer. Reference to polymerpreferably includes copolymers. “Copolymers” are polymers formed of morethan one polymer precursor. Polymers preferred for use in the drugformulation are those which are prepared from precursors that, in apreferred embodiment, are soluble in a solvent that is soluble in anantisolvent and can be polymerized with light initiation. One class ofsuch polymers includes those that are degradable, preferablybiodegradable. Another class of polymers includes poly lactic acids. Ina preferred embodiment, the polymers are degradable or erodible.Degradable or erodible polymers are those that degrade upon exposure tosome stimulus, including time and exposure to aqueous media. Degradableor erodible polymers include biodegradable polymers. Biodegradablepolymers degrade in a biological system, or under conditions present ina biological system such as an aqueous medium. Preferred biodegradablepolymers degrade within the device described herein following theintroduction of an aqueous medium. Examples of biodegradable polymersinclude those having at least some repeating unit representative of atleast one of the following: an alpha-hydroxycarboxylic acid, a cyclicdiester of an alpha-hydroxycarboxylic acid, a dioxanone, a lactone, acyclic carbonate, a cyclic oxalate, an epoxide, a glycol, andanhydrides. Preferred degradable or erodible polymers comprise at leastsome repeating units representative of polymerizing at least one oflactic acid, glycolic acid, lactide, glycolide, ethylene oxide andethylene glycol.

Another class of suitable polymers are biocompatible polymers. One typeof biocompatible polymers degrades to a nontoxic acidic product.Specific examples of biocompatible polymers that degrade to nontoxicproducts that do not need removal from biological systems includepoly(hydro acids), poly (L-lactic acid), poly (D,L-lactic acid), poly(glycolic acid) and copolymers thereof. Polyanhydrides have a history ofbiocompatibility and surface degradation characteristics (Langer, R.(1993) Acc. Chem. Res. 26:537-542; Brem, H. et al. (1995) Lancet345:1008-1012; Tamada, J. and Langer, R. J. (1992) J. Biomat Sci.-Polym.Ed. 3:315-353).

Poly(lactic-co-glycolic acid) (PLGA) is a copolymer prepared byco-polymerization of two different monomers, the cyclic dimers(1,4-dioxane-2,5-diones) of glycolic acid and lactic acid. Duringpolymerization, successive monomeric units of glycolic or lactic acidare linked together by ester linkages, thus yielding a linear, aliphaticpolyester as a product. Depending on the ratio of lactide to glycolideused for the polymerization, different forms of PLGA can be obtained.These forms are usually identified in regard to the monomers' ratioused, e.g. PLGA 75:25 identifies a copolymer whose composition is 75%lactic acid and 25% glycolic acid. PLGA degrades by hydrolysis of itsester linkages in the presence of water. The time required fordegradation of PLGA is related to the monomers' ratio used inproduction, where the higher the content of glycolide units, the lowerthe time required for degradation, other than a copolymer with a 50:50monomeric ratio of lactide and glycloide, which exhibits degradation inan aqueous environment in about two months. PLGA copolymers that areend-capped with esters (as opposed to the free carboxylic acid)demonstrate longer degradation half-lives. PLGA is a biodegradablepolymer because it undergoes hydrolysis in the body to produce theoriginal monomers, lactic acid and glycolic acid. These two monomersunder normal physiological conditions, are by-products of variousmetabolic pathways in the body. Since the body effectively deals withthe two monomers, there is very minimal systemic toxicity associatedwith using PLGA for drug delivery or biomaterial applications. Thepossibility to tailor the polymer degradation time by altering the ratioof the monomers used during synthesis lends this polymer particularlysuitable for the device herein, as further illustrated below (FIG. 11).A skilled artisan, however, will appreciate that other biodegradablepolymers are similarly beneficial for use as the solubility-modifyingexcipients, including polycaprolactone, polyglycolide, polylactic acid,and poly-3-hydroxybutyrate.

An exemplary approach for filling a drug delivery device reservoir witha drug formulation is now provided. The device, fitted at one end with astainless steel or sintered titanium porous partition or polymericmembrane, is weighed (including the end cap) to obtain a tear weight andpositioned in a holder with the partition end down. A known amount ofdrug, e.g., approximately 100 mg of risperidone (free base) powder(Ren-Pharm Intl, particle size 2-100 micrometers), is filled into eachreservoir. The powder may be filled by any known powder fillingapproaches and may be packed into the reservoir using a blunt 3.7 mmstainless steel rod. A drug/solubility-modifying agent mixture couldalso be introduced as a solution in a volatile solvent such as ethanol.Incremental introduction of the solution into the reservoir withconcomitant removal of the solvent, by for example elevated temperature,a stream of nitrogen, or reduced pressure, is another approach. A solidor semi-solid mixture of drug and solubility-modifying agent can also beextruded into the reservoir by standard methods known in the art.Following filling with dry drug, the open end of the reservoir is sealedusing a mechanically pressed-on end-cap described above and thereservoir reweighed and the drug fill weight calculated. Alternatively,the open end can be fitted with a second porous membrane to increase(double) the effective surface area available for diffuse release of thedrug, as illustrated in FIGS. 4-5.

The aqueous suspension, in one embodiment, is made by mixing acrystalline, amorphous, freeze-dried, spray dried or otherwise driedpowder, pellets or micronized powder of the agent (or the agent co-mixedas a physical mixture, melt or fusion mixture or with various bulkingagents and excipients known in the art), and solubility-modifyingexcipients with the internal aqueous medium under conditions designed toachieve the desired concentration gradient. A small portion of the dryagent dissolves in the aqueous solution at once, while the bulk remainsin insoluble suspension form. Initially, only the small proportion ofthe total amount of agent filled into the internal reservoir thatdissolves in the internal aqueous phase is available to diffuse out ofthe chamber through the porous partition. As the soluble form of theagent diffuses out of the reservoir through the partition into theexternal medium, it is replaced by the dissolution of the insoluble formwithin the reservoir. The balance between the exit of the initiallydissolved agent and the continuous process of dissolution of theinsoluble agent maintains a nearly constant concentration of solubleagent within the device reservoir.

At small scale production, the membrane-end of each reservoir isinserted into a tube connected to one end of a 3-way valve. Theconnection is sealed using an o-ring selected to fit around thereservoir housing and screw-cap compression fitting. A syringe filledwith phosphate buffered saline (PBS, pH 7.4) is attached to a secondport of the three-way valve. The third port is connected to a vacuumsource. Initially, the valve is positioned so that the reservoir isconnected to the vacuum source and the air in the reservoir is evacuatedfor 10 minutes. The valve is then positioned so that the evacuated, dry,drug-loaded reservoir is connected to the syringe containing buffer. PBSis drawn into the evacuated reservoir under reduced pressure createdduring the vacuum step, instantly creating a drug suspension in situwithin the reservoir. A removable cap is fixed to the membrane end ofthe reservoir until used.

For single devices fitted with membranes on each end, the device isfirst placed in a cylindrical tube with a slightly larger inner diameterthan the outer diameter of the device and a slightly longer overalllength. The tube is capped and sealed on one end. The open end of thetube is then connected to the to the 3-way valve, evacuated andback-filled with buffer as described above. Alternatively multipledevices may be placed in a larger chamber or manifold system andsimilarly filled with buffer following evacuation of the chamber undervacuum.

In larger scale production, individual drug-loaded, sealed reservoirsare placed in 13 mm neck borosilicate glass lyophilization vialsdesigned to hold each device. The vials are loosely fitted with alyophilization stopper and placed in a vacuum oven equipped withmechanically movable shelves. The vials are heated at 140° C. for 30minutes to terminally sterilize the systems. After return to ambienttemperature, a vacuum is applied for 24 hours at which time the vialsare sealed in vacuo by raising the shelf and fully seating the stoppers.Stoppers are affixed to the vials with removable aluminum crimps. Vialsare inspected, labeled and stored for use at room temperature.

B. Drug Delivery Device Performance: In Vitro Release

In one embodiment, the device is configured for, operable to, and/orcapable of delivering a sparingly soluble therapeutic agent at aconstant release rate for a period of at least about 1, 2, 3, 4, 5, 6,7, 8, 9, 10, 11, or 12 months, preferably, in one embodiment, at azero-order release rate. In vitro and in vivo performance of devices,prepared as described above, and capable of such delivery is nowdescribed.

As described in Example 1, typical Franz-type diffusion cells werefitted with a porous partition separating a donor chamber from areceptor chamber. The partitions were selected to span a wide range ofpore sizes (i.e., molecular weight cut-off or “MWCO”, values or nominalpose diameter) The porous partitions selected for use in this study hadmolecular weight cut-off values of 100,000 Daltons and 3,000 Daltons. Inaddition, a porous partition in the form of a 0.45 μmpolyvinyldenefluoride membrane filter (Durapore®) was used. A drugformulation in the form of an aqueous suspension comprising risperidone,as a model for a sparing soluble therapeutic agent, asolubility-modifying excipient PLGA (polylactic ploy glycolic acidco-polymer), a biodegradable, biocompatible copolymer, was prepared andplaced in the donor chamber. Release of risperidone from the formulationinto the receptor chamber was measured as a function of time, and theresults are shown in FIGS. 7A-7B.

FIGS. 7A-7B are graphs showing cumulative release of risperidone invitro, in mg (FIG. 7A), and the release rate of risperidone, in mg/day(FIG. 7B), as a function of time, in days. Sustained, substantiallyzero-order release rate was obtained, as best seen in FIG. 7B.Risperidone was released from the devices with the porous partitions of100,000 Dalton molecular weight cut-off (triangles), 3,000 Daltonmolecular weight cut-off (squares), and the 0.45 μmpolyvinyldenefluoride membrane filter (circles) at rates of 200 μg/day,100 μg/day and 70 μg/day, respectively, were achieved following a short3-4 day equilibration phase.

The aqueous solubility of many sparingly soluble drugs, such asrisperidone, increases as the pH is lowered to below that characteristicof physiological fluids (i.e., pH of 7.4). Risperidone is only sparinglysoluble in water at pH 7.4 (˜0.42 mg/mL). Assuming a risperidoneconcentration of 0.42 mg/mL, a pore area (A) of a 3.5 mm partition is0.09 cm², that two such partitions are used at each end of a drugdelivery device as described herein to provide a total diffusional areaof 2×A, that a diffusion constant of glucose is 7.0×10⁻⁶, and that thethickness of the partition is 0.12 cm, Fick's law can be used todetermine the rate of efflux, J, as follows:

$\begin{matrix}{J = {A\left( {D \cdot \frac{C}{T}} \right)}} & (5) \\{J = {0.01\mspace{14mu} {{cm}^{2}\left( {7 \times 10^{- 6}{cm}^{2}\sec^{- 1} \times \frac{0.42\mspace{14mu} {{mg} \cdot {ml}^{- 1}}}{0.12\mspace{14mu} {cm}}} \right)}}} & (6) \\{J = {0.012\mspace{14mu} {{mg} \cdot {day}^{- 1}}}} & (7)\end{matrix}$

This rate of efflux is well below the risperidone dose needed to providehuman therapeutic effects. An in vitro rate of about 1-2 mg/day would beneeded to achieve a therapeutic effect. To adjust the solubility ofrisperidone to a concentration within the device that would provide anefflux rate of between about 1-2 mg/day it would be necessary toincrease the internal concentration of soluble risperidone to 1-2 mg/mL.Therefore, solubility enhancement of the sparingly soluble drug isneeded to achieve a release rate from the device that provides atherapeutic dose. One approach is to lower the pH of the aqueoussuspension in the device. For example, at pH 6.8 the equilibriumconcentration of risperidone is about 1.5 mg/mL (see FIG. 6). At thisconcentration in the example given above the output rate would be intherapeutic range. But, since the device reservoir containing theaqueous suspension of drug communicates with the external medium in theenvironment of use that is filled with interstitial fluid, the pH ofwhich is regulated by buffers (primarily bicarbonate-pCO₂) to be 7.4, itis inevitable that the internal pH of the device would equilibrate to pH7.4 quickly following implantation, regardless of the initial pH of theinternal aqueous phase. The present device is based on a discovery thatcircumvents this dilemma. Inclusion of a solubility-modifying excipientscapable of generating acid groups for a sustained period of timemaintains a pH of the aqueous suspension in the device that provides asolubility of the drug sufficient to achieve a therapeutic rate ofrelease from the device for a sustained period of time. In oneembodiment, the solubility-modifying excipients is a biocompatible,biodegradable co-polymer of poly-lactic and poly-glycolic acids (PLGA).The PLGA effectively lowers the internal pH of the aqueous suspension,as the polymer erodes and releases acid groups, free lactic and glycolicacids, respectively. The continual release of acid groups increases thesolubility of the drug in the aqueous suspension, and provides increasedrelease rates despite the pH of external buffer being maintained at pH7.4.

This effect is illustrated by the study described in Example 3, whererelease of the sparingly soluble drug risperiodone, as a model of allsparingly soluble drugs, from aqueous suspensions comprising variousamounts of PLGA was measured as a function of time. As seen in FIG. 8,addition of 5 mg (diamonds), 10 mg (squares) or 15 mg (triangles) of85:15 PLGA to the aqueous risperidone suspension contained in thereservoir of the devices provided a dose-dependent increase in drugrelease rate. Drug formulations with 15 mg PLGA (triangles) yielded arelease rate that approached about 0.6 to 1.0 mg/day. Drug deliverydevices with increasing amounts of solubility-modifying excipient PLGAfrom 5 mg, 10 mg to 15 mg, gave a linear increase in release rate, from0.25 mg/day to 0.43 mg/day to 0.64 mg/day, as seen in FIG. 9. Doublingthe amount of solubility-modifying excipient PLGA doubled the releaserate.

Further studies were done to evaluate the type and combinations ofsolubility-modifying excipients. As discussed above, PLGA is prepared byco-polymerization of two different monomers, glycolic acid and lacticacid. The ratio of lactide to glycolide used in the synthesis yieldsdifferent forms of PLGA. A study was conducted to evaluate the releaserate of a sparingly-soluble drug from devices as described herein usingvarious types of PLGA. As described in Example 3, aqueous suspensions ofrisperidone five different PLGA co-polymers were prepared. The threeco-polymers were 50:50 poly(dl-lactide-co-glycolide-1A), 50:50poly(dl-lactide-co-glycolide-2A) and 85:15poly(dl-lactide-co-glycolide). Results are shown in FIG. 11.

FIG. 11 shows the risperidone release rate, in mg/day, as a function oftime, in days, for devices filled with the various drug formulations.The series of curves illustrate the influence of thesolubility-modifying agent PLGA with different lactic acid:glycoloicacid ratios on the in vitro release of risperidone. The first two curvesin FIG. 11 show the release of risperidone from an aqueous suspensioncomprising a 50:50 PLGA (DL1A and DL2A). The third curve shows therelease of risperidine from an aqueous suspension comprising 85:15 PLGA.The fourth and fifth curves illustrate the projected release ofrisperidone from an aqueous suspension comprising PLGA with ratios of90:10 and 95:5, respectively. The data and projections presented in FIG.11 illustrate that the erosion rate of PLGAs in aqueous media decreaseswith increasing lactic acid content. Aqueous mixtures of PLGAs ofincreasing lactic acid content (starting at 50:50) can be selected togenerate a constant number of free acid equivalents (lactic acidglycolic acid) over a selected period of time, for example for a periodof 2, 3, 4, 5, 6 months or longer. The free lactic acid and glycolicacid equivalents generated by the differential erosion of the variousco-polymers in the PLGA co-polymer family increase the solubility ofrisperidone and thus increase its diffusion rate across the porouspartition of the drug delivery device. By selecting mixtures of PLGAswith different, and in one embodiment, overlapping erosion rates, theaugmented, free acid-mediated release of a sparingly soluble drug ismaintained for a sustained period. In practice, drug delivery devicescan be filled with dry powder physical mixtures of risperidone and oneor more PLGAs, selected to provide sustained release of the sparinglysoluble drug. Following hydration and implantation the PLGAs begin toerode at a rate depending on its ratio of lactic:glycolic acid,continuously generating free acid equivalents which, in turn, elevatethe solubility of the sparingly soluble drug within the reservoir andthus sustain its constant release rate from the device. Release rates inthe therapeutic range of 1-2 mg/day are achievable for periods rangingfrom of 1-6 months.

Accordingly, in one embodiment, devices comprising a drug formulationwith a solubility-modifying agent are contemplated, wherein thesolubility-modifying agent is one or more poly(dl-lactide-co-glycolide)co-polymers, selected to have a glycolide content of less than 15%,preferably less than 10%, more preferably less than 8%, 7%, 6%, 5%, 4%,3%, 2% or 1%, to achieve a sustained, continuous release of acid groupsin the drug formulation, thereby achieving a sustained, zero-orderrelease of a sparingly soluble drug. In one embodiment, the drugformulation is an aqueous suspension comprising the sparingly-solubledrug. That is, the reservoir of the device prior to use comprises anaqueous suspension of the drug and one or more solubility-modifyingexcipients. In another embodiment, the drug formulation is a dry ordried mixture of the sparingly-soluble drug and the one or moresolubility-modifying excipients. In this latter embodiment, the dry drugformulation hydrates subsequent to implantation of the device by in-flowof interstitial fluid across the porous partition.

In another study, described in Example 4, devices were prepared fromtitanium cylindrical tubes. The reservoir of each device was filled withrisperidone free base in the form of a dry powder, and in one group ofthe devices, a mixture of 50:50 PLGA and 85:15 PLGA was included. Thedevices were fitted with a porous partition at each end, as illustratedin FIGS. 4A-4B and 5A-5B. Release of risperidone from the devices into arecipient buffer was measured and the results are shown in FIG. 11. Asseen, without the added PLGA the devices released risperadone at a rateof less than 0.05 mg/day, far below the therapeutic dose level of >1gm/day. In contrast, the devices with the added PLGA releasedrisperidone at an average of 1-1.2 mg/day, well within the targettherapeutic range.

In yet another study, devices were prepared as described herein antested in vitro and in vivo. As described in Example 5, devices having acylindrical titanium housing with an internal reservoir were filled withdry powdered risperidone. One group of the devices additionally includeda mixture of 50:50 PLGA and 85:15 PLGA admixed with the risperidone. Asubset of each group of devices was tested for in vitro release ofrisperidone and the other subset implanted subcutaneously into rats.Release of risperidone was determine by taking aliquots of buffer orblood, accordingly. Results are shown in FIGS. 12A-12B.

FIG. 12A shows in vitro release rates for the devices containing onlyrisperidone (triangles) compared to devices containing risperidone plusPLGA (a mixture of PLGA 50:50 and PLGA 85:15, squares). The deviceswithout PLGA shows limited release; devices with PLGA maintain a releaserate above 0.4 mg/day for 30 days while those without PLGA release lessthan 0.05 mg/day throughout the same period.

FIG. 12B presents the comparable in vivo data. For the group containingrisperidone only (triangles), plasma levels of risperidone (plus9-hydroxyrisperidone) are barely detectable for the first week and thenfall below the limits of quantitation of the assay (0.1 ng/mL). In thedevices including a solubility-modifying agent (squares), plasma levelsof risperidone (plus 9-hydroxyrisperidone) quickly build and maintain alevel above 4 ng/mL for the entire one month period.

An excellent in vitro/in vivo correlation is apparent by comparing thedata presented in FIGS. 12A-12B. It is apparent from this data that itis possible to optimize the in vivo performance of the devices using invitro data. Increasing or prolonging output (or both) is possible byselection of the appropriate mass and erosion rates of thesolubility-modifying agent(s).

Although the specific examples given herein are with respect to themodel drug risperidone, it will be appreciated that any sparing solubledrug can be used in the device described herein. Example 6 illustratesdevices designed for delivery of asenapine, a drug with a watersolubility of 3 mg/mL at pH 7-7.4. The table below presents thesolubility and pharmacokinetic parameters of exemplary drugscontemplated for use in the devices described herein.

Olanzapine Paliperidone Aripiprazole Asenapine Risperidone HaloperidolParameter CL(L/hr) (iv) 21 2.37 52 2.84 42; 22 CL (L/day) 504 57 1248 68960 Target Steady state 30 100 2.1 20 2 plasma conc (ng/mL) Output rate15 5.7 2.62 1.36 1.92 Half-life (hr) 45 24 3 24 3 29 Bioavailabilityoral 60% 28% 87%  2% 65 65% sublingual 30% Water solubility (mg/mL)neutral pH 0.54 3.0 0.48 0.025 acidic pH 100 × neutral 13 4.2

III. Methods of Treatment

In another aspect, methods for treating diseases or disorders using thedevice described herein above are provided. In particular embodiments,the implantable device delivers a therapeutic agent for between 1-3months, more preferably for between 1-4 months, still more preferablyfor between 1-5 months, 1-6 months, 2-4 months, 2-6 months, or between2-12 months. For chronic diseases, such as schizophrenia, bipolardisorder or alcoholism, patients may be treated from many months,perhaps years. In such disease settings, when one device is expendedafter several months of operation, it would be removed and replaced witha new, fully-charged device in order to provide uninterrupted therapy.The device can be implanted at the same, or a different site.

IV. Examples

The following examples are illustrative in nature and are in no wayintended to be limiting.

Example 1 In Vitro Release of Risperidone

Typical Franz-type diffusion cells were fitted with porous partitionsseparating a donor chamber from a receptor chamber. The porouspartitions were membranes with pore sizes of 3,000 Daltons molecularweight cut-off, 100,000 Daltons molecular weight cut-off and a 0.45micron polyvinyldenefluoride membrane filter (Durapore®). An aqueoussuspension comprising risperidone (10 mg/mL PBS, pH 7.4) andpoly(lactic-co-glycolic acid) was prepared and placed in the donorchamber of each cell. The cells were maintained at room temperature.Risperidone concentration in the buffer in the receptor chamber wasmeasured by UV spectroscopy at 270 nm using an extinction coefficientfor risperidone of 20.5. Buffer was diluted to 20-50 micrograms/mL priorto measurement to insure that the concentration was within the linearrange of the UV measurement. Buffer was replaced each day. Results areshown in FIGS. 7A-7B.

Example 2 In Vitro Release of Risperidone

Franz diffusion cells were prepared as noted above. Aqueous suspensionscomprising risperidone (10 mg) and 5 mg, 10 mg, or 15 mg of 85:15poly(lactic-co-glycolic acid) were prepared and placed in the donorchamber of each cell. The cells were maintained at room temperature.Risperidone concentration in the buffer in the receptor chamber wasmeasured by UV spectroscopy at 270 nm using an extinction coefficientfor risperidone of 20.5. Buffer was diluted to 20-50 micrograms/mL priorto measurement to insure that the concentration was within the linearrange of the UV measurement. Buffer was replaced each day. Results areshown in FIG. 8.

Example 3 In Vitro Release of Risperidone

Two groups of drug delivery devices were assembled as follows.Cylindrical titanium tubes, 4 mm in diameter and 35 mm in length, wereprepared. Risperidone free base in the form of a dry powder (100 mg,Ren-Pharm Intl, particle size 2-100 micrometers) was filled into the afirst group of devices (Group A, n=9). In a second group of devices(Group B, n=5), the same amount of risperidone was mixed with 5 mg 50:50PLGA and with 15 mg 85:15 PLGA (Lakeshore Biomaterials, Birmingham,Ala.). The reservoirs of the devices were filled with phosphate bufferedsaline (PBS) at pH 7.4 under reduced pressure. Each of the devices inGroup B was fitted with two regenerated cellulose membranes (5000Daltons molecular weight cut off) at each end of the cylindrical tube bysandwiching the membrane between an o-ring and titanium frit (5 micronpores), as illustrated in FIGS. 5A-5B. The devices in Group A wereidentical, except the cellulose membrane was omitted, as illustrated inFIGS. 4A-4B. The devices were then placed individually in tubescontaining approximately 1 mL PBS. The tubes were placed on a rotatingrack (1 revolution per hour) at 37° C.

Risperidone concentration in the recipient buffer was measured by UVspectroscopy at 270 nm using an extinction coefficient for risperidoneof 20.5. Buffer was diluted to 20-50 micrograms/mL prior to measurementto insure that the concentration was within the linear range of the UVmeasurement. Buffer was replaced each day. The release rates for thedevices in Group B (closed circles, with two different PLGAs) and GroupA (open circles, without PLGA) are shown in FIG. 10. Devices without asolubility-modifying excipients released at a rate of less than 0.05mg/day, far below the therapeutic dose level of >1 gm/day. In contrast,the devices with a solubility-modifying excipient, a mixture of twodifferent PLGAs, released with an average of 1-1.2 mg/day, well withinthe target therapeutic range for risperidone.

Example 4 Devices with Mixture of PLGAs with Differential AqueousErosion Rates

The influence of lactic acid:glycolic acid ratio of PLGAs on the invitro release rate of risperidone was investigated. Physical mixtures of100 mg of risperidone and 5-15 mg of PLGA with lactic acid:glycoloicacid ratios of 50:50 and 85:15 were made and suspended in phosphatebuffered saline. Risperidone release was measured as described inExample 1. Results are shown in FIG. 11. The first two curves in FIG. 11show the release of risperidone mixed with two different samples of a50:50 PLGA (DL1A and DL2A supplied by Lakeshore Biomaterials,Birmingham, Ala.). The third curve shows the release of risperidine whenmixed with 15 mg of 85:15 PLGA (also obtained from LakeshoreBiomaterials, Birmingham, Ala.). The fourth and fifth curves illustratethe projected release of risperidone when mixed with PLGA with ratios of90:10 and 95:5, respectively.

Example 5 In Vitro and In Vivo Delivery

Two groups of devices (n=6) were prepared. The main body of each deviceconsisted of a cylindrical titanium tube, definining an internalreservoir. The tubes were 35 mm in length, with an outer diameter (OD)of 4.3 mm and an inner diameter (ID) of about 3 mm (PeridotManufacturing, Pleasanton, Calif.). Both ends of each tube were tapered(lathed) so as to accept a press cap which served to retain a porouspartition, as illustrated in FIGS. 4-5. The first group of devices wasfitted on one end with a titanium sintered frit having a diameter 3.5mm, a depth 0.5 mm, and a 5 micron nominal particle retention size(Applied Porous Materials, Farmington, Conn.). The frit was secured ineach tube by placing the tube horizontally in a holder, and thenpositioning over the open end, in sequence, the frit, an o-ring and thepress cap. The frit/o-ring stack was sealed in place on the edges of thereservoir by pressing downward the on the membrane press cup (PeridotManufacturing, Pleasanton, Calif.) using a mechanical press (SchmidtPress No 307).

The devices were inverted and filled with 100 mg of dry powderedrisperidone (Ren-Pharm Intl, particle size 2-100 micrometers). A 2.5 mmstainless steel ball was placed in the reservoir to provide mixing. Asecond frit was sealed to the remaining open end of the reservoir usingthe same procedure as described above for the other end.

The second group of devices were each fitted according to the sameprocedure, but include a circular 3.5 mm diameter polymer membrane (3KDMWCO Membranes, Millipore #PLBCO2510) positioned between the frit ano-ring as illustrated in FIGS. 5A-5B. The polymer membrane was intendedto serve as a biocompatible interface between the frit and thesubcutaneous tissue after implantation. The devices in the second groupwere filled with a physical mixture composed of 100 mg risperidone, 15mg 50:50 PLGA and 50 mg 85:15 PLGA (Lakeshore Biomaterials, Birmingham,Ala.), plus the 2.5 mm stainless steel mixing ball.

After filling with drug and/or PLGA, all of the devices in each groupwere hydrated by placing each into a vacuum chamber, evacuating thechamber under vacuo, and back-filling with degassed PBS.

Both groups of devices were equally split, each subgroup of 3 devicesdestined for either in vitro or in vivo testing. In vitro release ratewas determined by submerging the devices in 1 mL PBS held in screw capcryo-vials. The tubes were sealed and placed in a rotating rack (1revolution per hour) in a 37° C. incubator. Each day, each device wasmoved to a new tube containing fresh buffer, and the risperidoneconcentration in the release buffer was measured by UV spectroscopy.

For in vivo experiments, rats were anesthetized and devices placedsubcutaneously in the dorsum, lateral to the midline, using a trocar.Whole blood samples were obtained by tail vein venous puncture. Plasmawas immediately separated from formed elements by centrifugation. Plasmasamples were sent to a reference laboratory (Integrated AnalyticsSolutions, Berkeley, Calif.) for detection of risperidone and9-hydroxyrisperidone concentrations in the plasma, and were quantifiedby a validated LC/MS/MS method. The lower limit of quantitation was 0.10ng/mL. Results are shown in FIGS. 12A-12B.

FIG. 12A shows in vitro release rates for the devices containing onlyrisperidone compared to devices containing risperidone plus PLGA (amixture of PLGA 50:50 and PLGA 85:15 as described above). The groupwithout PLGA shows limited release; devices with PLGA maintain a releaserate above 0.4 mg/day for 30 days while those without PLGA release lessthan 0.05 mg/day throughout the same period.

FIG. 12B presents the comparable in vivo data. For the group containingrisperidone only, plasma levels of risperidone (plus9-hydroxyrisperidone) are barely detectable for the first week and thenfall below the limits of quantitation of the assay (0.1 ng/mL). In thegroup with PLGA, on the other hand, plasma levels of risperidone (plus9-hydroxyrisperidone) quickly build and maintain a level above 4 ng/mLfor the entire one month period. An excellent in vitro/in vivocorrelation is apparent by comparing the data presented in FIGS. 12A and12B.

Example 6 Device for Delivery of Asenapine

Asenapine exemplifies an agent that will benefit from the devicedescribed herein. This agent has recognized activity in the treatment ofschizophrenia and bipolar disorder but suffers from a number of factorsthat detract from its clinical benefit. It has poor oral bioavailability(<2%) due to a high first pass effect; its current dosage form is asublingual tablet that must be taken twice daily, making compliance inthe target patient population challenging. Asenapine is cleared rapidlyfrom circulation (plasma clearance 52 L/hr) but is active at low plasmaconcentrations (70% D₂ receptor occupancy at 2.1 ng/mL). For the drugdelivery device enabled herein, the required output rate can becalculated by multiplying the CL (in mL/day) times the steady stateplasma concentration required for efficacy (mg/mL): =1.25×10⁶mL/day×2.1×10⁻⁶ mg/mL=2.65 mg/day output to maintain 2.1 ng/mL.

The drug exhibits limited water solubility at neutral pH (3 mg/mL at pH7.0) but its solubility improves as the pH is lowered (13 mg/mL in 0.1NHCl).

The drug delivery device claimed herein typically would consist of atitanium cylinder of 3.8 mm ID and 40 mm in length fitted at each endwith sintered titanium frits of 0.2 cm in thickness and a pore size of≦2 μm. The volume of the reservoir=πr²h=453 The combined surface area ofthe two frits is 0.23 cm²; the porosity of the frits is 30%, so theeffective pore area available for diffusion is 0.07 cm². If thereservoir were filled with an aqueous suspension of 250 mg of just thedrug, the flux (the outward diffusion) at neutral pH can be calculatedfrom Fick's law: J(mg/sec)=D(cm/sec)×ΔC(the concentration gradientbetween the inside and outside of the reservoir expressed as mg/cm³)/0.2cm (the thickness of the diffusion membrane, or frit thickness in thiscase)×A (effective pore area available for diffusion expressed as cm²).The concentration in the external medium (intersitial fluid) can beconsidered to be zero. The concentration within the aqueous phase of thereservoir will be 3 mg/mL, the drug's inherent solubility at neutral pH,so ΔC=3 mg/mL, or 3 mg/cm³. The diffusion coefficient of the drug can beapproximated to be that of a small molecule such as glucose (7×10⁻⁶cm²/sec). Thus at neutral pH, flux would be 7×10⁻⁶ cm·sec⁻¹×3mg·cm³⁻¹/0.2 cm×0.07 cm²=7.35×10⁻⁶ mg·sec⁻¹. Converting J to mg/day, themaximum flux under these conditions would be 0.6 mg/day. This levelfalls short of the output needed to provide a therapeutic dose of thedrug, 2.65 mg/day, by fourfold.

The current device provides a means to boost the output rate ofasenapine by loading a mixture of PLGA polymers together with the drug.In this case 250 mg of drug plus approximately 200 mg of PLGA would beloaded into the 450 μL reservoir. The device is hydrated and implantedunder the skin. As the PLGA hydrolyses over time, lactic acid andglycolic acid equivalents are generated from the insoluble polymerwithin the reservoir, effectively lowering the internal pH. Furthermore,the lactide and glycolide monomers form a salt with the asenapine baseand improve the drug's water solubility within the reservoir to at least13 mg/mL. Recalculating flux under these conditions the output ratewould be 2.6 mg/day, and, given the total drug loaded is 250 mg, thisoutput can be maintained for 96 days, thus providing a therapeutic levelof drug for a 3-month period following implantation.

Comparative Example 1

Prior art devices as described in U.S. Pat. Nos. 3,896,819; 3,948,254;3,948,262; 3,993,072; and 3,993,073 were examined and compared to thepresent device. The deficiency of the prior art devices is apparent byexamining the in vivo data presented in FIG. 12B, wherein two devicesare compared; one modeled after the teachings of these prior artdocuments and the other after the present teachings. In this comparison,devices are cylindrical reservoirs, the outer surface of which is asmooth biocompatible, satin-finished surface, suitable for implantationunder the skin. The ends of the reservoir are fitted with porouspartitions. The bioavailability data for such devices containing onlyrisperidone (lower curve, triangles) and modeled after the prior artteaching (i.e., “surrounded by an enclosure that at least is partiallyformed by a microporous membrane designed to be permeable to the drug”)do not provide output sufficient to maintain a therapeutic dose. This isbecause the aqueous solubility of risperidone is too low to create aconcentration gradient sufficient to drive mass transport of the drugacross the porous surface of devices with this configuration. Incontrast, devices containing a preferred solubility-modifying excipientas described herein produce a therapeutic level active drug for at leastone month (upper curve, squares).

The limitations of the prior art device, in the case of risperidone, areillustrated in the following analysis. The aqueous solubility ofrisperidone at pH 7.4 is about 0.42 mg/mL. Thus, when the carrier fluidis water and an aqueous suspension of risperidone is loaded into thedevice contemplated in the prior art device, and applying Fick's law,the porous surface area required to produce an output rate of about 1.78mg/day (i.e., the dose required to maintain a therapeutic response ofthis drug) can be calculated.

The assumptions for this illustration are:

-   -   i. The equilibrium solubility of risperidone free base in PBS is        0.42 mg/mL at pH 7.4.    -   ii. The external concentration will effectively be zero, so the        concentration gradient is approximated by the internal        concentration (i.e., 0.42 mg/mL).    -   iii. The output rate (release rate) desired for the product,        based on the dose rate of Risperdal Consta, is 1.78 mg/day        (i.e., based on the recommended dose of Risperdal Consta of 25        mg dose every 2 weeks, www.risperdalconsta.com).    -   iv. The diffusion constant of risperidone is equal to that of        glucose (7×10⁻⁶ cm²/sec,        http://www.d.umn.edu/˜dlong/samrpt.html).    -   v. The thickness of porous portion of the encasement is 1.2×10⁻¹        cm.    -   vi. The porosity of the porous portion of the encasement is 50%.

If the OD of the porous encasement is 4.6 mm and the ID is 2.2 mm, thesurface area of the inner exposure of the porous segment needed toprovide the desired release rate is calculated as follows. First, therelease rate, is restated in terms of mg per sec:

$\begin{matrix}{{{Release}\mspace{14mu} {{Rate}(R)}} = {\frac{1.78\mspace{14mu} {mg}}{day} = {\frac{1.78\mspace{14mu} {mg}}{8.64 \times 10^{4}\sec} = {2.05 \times 10^{- 5}{{mg} \cdot \sec^{- 1}}}}}} & (8)\end{matrix}$

Flux is expressed in terms of the number of molecules (or mass) whichdiffuse through a defined area per unit time. In this case flux has theunits mg·sec⁻¹·cm³⁻¹, accounting for the movement of molecules through adefined surface area of the membrane (which shall be called “A”). It isthis area that we wish to determine.

Fick's law states (units in parentheses):

$\begin{matrix}{{J\left( {{mg} \cdot {cm}^{2^{- 1}} \cdot \sec^{- 1}} \right)} = {{D\left( {{cm}^{2} \cdot \sec^{- 1}} \right)} \cdot \frac{C\left( {{mg} \cdot {cm}^{3^{- 1}}} \right)}{T({cm})}}} & (9)\end{matrix}$

Expressing flux in terms of release rate (“R”), it's necessary toinclude the surface area available for diffusion, that is: R÷membranesurface area (A):

$\begin{matrix}{{C\left( {{mg} \cdot {cm}^{3^{- 1}}} \right)} = \frac{\frac{R\left( {{mg} \cdot \sec^{- 1}} \right)}{A\left( {cm}^{2} \right)} \cdot {T({cm})}}{D\left( {{cm}^{2} \cdot \sec^{- 1}} \right)}} & (10)\end{matrix}$

Wherein D is the diffusion constant of the agent; and T_(M)=thethickness of the porous section. Inserting the values for concentration,release rate and thickness:

$\begin{matrix}{{0.4\mspace{14mu} {{mg} \cdot {cm}^{3^{- 1}}}} = \frac{\frac{2.05 \times 10^{- 5}{{mg} \cdot \sec^{- 1}}}{A} \times 1.2 \times 10^{- 1}{cm}}{7 \times {10^{- 6} \cdot {cm}^{2} \cdot \sec^{- 1}}}} & (11)\end{matrix}$

Rearranging and solving for A:

$\begin{matrix}{{A = \frac{2.47 \times 10^{- 6}{{mg} \cdot {cm} \cdot \sec^{- 1}}}{\left( {7 \times {10^{- 6} \cdot {cm}^{2} \cdot \sec^{- 1}}} \right) \cdot \left( {0.42\mspace{14mu} {{mg} \cdot {cm}^{3^{- 1}}}} \right)}}{A = {0.84\mspace{14mu} {cm}^{2}}}} & (12)\end{matrix}$

In practice, only up to about 50% of this area is available fordiffusion, so the effective porous surface area (A_(Eff)) needed is:

$\begin{matrix}{A_{Eff} = {\frac{0.84\mspace{14mu} {cm}^{2}}{.50} = {1.68\mspace{14mu} {cm}^{2}}}} & (13)\end{matrix}$

Thus, in the case of risperidone, owing to its low aqueous solubility, adevice modeled after the prior art device teachings would require atleast 1.7 cm² of porous surface area in order to provide a therapeuticdose in vivo.

For the reasons detailed above, the device as described herein, which isdesigned to be implanted subcutaneously, is cylindrical with smoothwalls. The porous partition needed for drug diffusion is fitted on oneor both ends of the cylindrical reservoir. A typical device would havean OD of 0.46 cm so that the device can be comfortably placed under theskin. Given the thickness of the wall required for mechanical strength,the ID of the reservoir is 0.22 cm. So the maximum porous surface areaavailable for a device with smooth tubular surfaces and fitted at bothends with a circular porous partition is:

A=πr ²=π(0.11)²=0.04 cm²  (14)

and for two membranes

0.04 cm²×2=0.08 cm²  (15)

A device of this size and shape, which is suitable for subcutaneousimplantation, using an aqueous carrier fluid and modeled on theteachings of the prior art device, would not provide sufficient drugoutput when risperidone or other agents with similarly low aqueoussolubility are selected. As shown in the calculations above, followingthe teachings of the prior art device, an effective porous surface areaof at least 1.7 cm² would be needed to provide a therapeutic dose ofrisperidone (Equation 13). However, in the configuration describedherein which is suitable for subcutaneous implantation, the porouspartition provides only a 0.08 cm² of effective surface area, 20 timesbelow the value needed by the prior art device.

As described below, and presented graphically in FIG. 12A-12B, thecurrent device overcomes the limitations of both the prior art devicesand achieves a sufficient output rate of risperidone and other drugswith similarly low aqueous solubility properties.

While a number of exemplary aspects and embodiments have been discussedabove, those of skill in the art will recognize certain modifications,permutations, additions and sub-combinations thereof. It is thereforeintended that the following appended claims and claims hereafterintroduced are interpreted to include all such modifications,permutations, additions and sub-combinations as are within their truespirit and scope.

It is claimed:
 1. A method for delivering a sparingly water soluble drugfrom an aqueous suspension into an environment of use, comprising:providing a device comprising a porous partition with pores having apore size and with a porosity, a reservoir and a formulation containedin the reservoir, the formulation comprising a sparingly water solubledrug having a soluble form and an insoluble form and asolubility-modifying excipient that generates acidic groups for a periodof between about 2-12 months to provide, when the formulation ishydrated to form an aqueous suspension, a concentration of drug in thesoluble form that is released across the porous partition to provide atherapeutic dose of the drug for the period, wherein release of thetherapeutic dose is dependent on pH of the drug formulation andindependent of pore size of the porous partition with said porosity. 2.The method according to claim 1, wherein the excipient is abiocompatible, bioerodible polymer.
 3. The method according to claim 2,wherein the polymer is selected from the group of polylactides,polyglycolides, copolymers thereof, and polyethyleneglycol.
 4. Themethod according to claim 2, wherein the polymer is a co-polymer ofpolylactic acid and polyglycolic acid monomeric units, wherein thepolylactic acid content is between about 50% to 100%.
 5. The methodaccording to claim 1, wherein the solubility-modifying excipient iswater insoluble.
 6. The method according to claim 5, wherein thesolubility-modifying excipient is retained by the porous partition. 7.The method according to claim 1, wherein the solubility-modifyingexcipient is a water insoluble polymer.
 8. The method according to claim7, wherein the solubility-modifying excipient undergoes hydrolysis toform a water-soluble monomer.
 9. The method according to claim 1,wherein the solubility-modifying excipient is an erodible polymer or adegradable polymer.
 10. The method according to claim 9, wherein thepolymer is polymerized from a monomer selected from the group consistingof lactic acid, glycolic acid, lactide, glycolide, ethylene oxide andethylene glycol.
 11. The method according to claim 9, wherein thepolymer is a poly(lactic-co-glycolic acid) copolymer.
 12. The methodaccording to claim 1, wherein said drug is a neuroleptic agent.
 13. Themethod according to claim 12, wherein the neuroleptic agent isrisperidone, 9-hydroxyrisperidone, olanzapine, or a pharmaceuticallyacceptable salt thereof.
 14. The method according to claim 1, whereinsaid drug is risperidone.
 15. The method according to claim 1, whereinthe drug is selected from the group consisting of buprenorphine,paliperidone, asenapine, haloperidol, aripiprazole, and pharmaceuticallyacceptable salts thereof.
 16. A method for treating a patient with asparingly water soluble drug, comprising: implanting a device comprisinga porous partition with pores having a pore size and with a porosity, areservoir and a formulation contained in the reservoir, the formulationcomprising a sparingly water soluble drug having a soluble form and aninsoluble form and a solubility-modifying excipient that generatesacidic groups for a period of between about 2-12 months to provide, whenthe formulation is hydrated to form an aqueous suspension, aconcentration of drug in the soluble form that is released across theporous partition to provide a therapeutic dose of the drug for theperiod, wherein release of the therapeutic dose is dependent on pH ofthe drug formulation and independent of pore size of the porouspartition with said porosity.
 17. The method according to claim 16,wherein the device is implanted subcutaneously.
 18. A drug deliverydevice, comprising: a non-erodible, non-porous housing member defining areservoir; a porous partition positioned for communication with thereservoir, the porous partition having a porosity and pores with a poresize; contained within said reservoir, a drug formulation comprising asparingly water soluble drug having a soluble form and an insoluble formand a solubility-modifying excipient that generates acidic groups for aperiod of between about 2-12 months to provide, when the formulation ishydrated to form an aqueous suspension, a concentration of drug in thesoluble form that is released across the porous partition to provide atherapeutic dose of the drug for the period, wherein release of thetherapeutic dose is dependent on pH of the drug formulation andindependent of pore size of the porous partition with said porosity.